The implantation of stents has become established as one of the most effective therapeutic measures for the treatment of vascular diseases. Stents have the purpose of assuming a supporting function in hollow organs of a patient. For this purpose, stents featuring conventional designs have a filigree supporting structure comprising metal struts, which is initially present in compressed form for introduction into the body and is expanded at the site of the application. One of the main fields of application of such stents is to permanently or temporarily widen and hold open vascular constrictions, particularly constrictions (stenosis) of coronary vessels. In addition, aneurysm stents are known, which are used to stabilize damaged vessel walls.
Stents have a tubular base body, through which the blood continues to flow without impairment and the circumferential wall of which performs a supporting function for the vessel wall. The base body is frequently implemented as a mesh-like structure, having a plurality of individual strut sections that are connected to each other. The base body of the stent is made of an implant material. An implant material is a non-living material, which is used for applications in medicine and interacts with biological systems. A basic prerequisite for the use of a material as an implant material, which is in contact with the surrounding body area when used as intended, is the body friendliness thereof (biocompatibility). Biocompatibility shall be understood as the ability of a material to evoke an appropriate tissue response in a specific application. This includes an adaptation of the chemical, physical, biological, and morphological surface properties of an implant to the recipient's tissue with the aim of a clinically desired interaction. The biocompatibility of the implant material is also dependent on the temporal course of the response of the biosystem in which it is implanted. For example, irritations and inflammations can occur in a relatively short time, which can lead to tissue changes. Depending on the properties of the implant material, biological systems thus react in different ways. According to the response of the biosystem, the implant materials can be divided into bioactive, bioinert and degradable/resorbable materials.
Only metal implant materials for stents are of interest for the purpose of the present invention, and more particularly biocorrodible alloys of the element magnesium, containing zinc and/or aluminum as the main minor element. Such implants made of biocorrodible alloys can also be coated with biocompatible polymers.
As was already mentioned above, in addition to fulfilling the desired mechanical properties, a stent should be made of a biocompatible material so as to minimize rejection reactions. For example, stents are used in approximately 70% of all percutaneous interventions, however in a significant number of cases in-stent restenosis occurs due to excessive neointimal growth, which is caused by a strong proliferation of the arterial smooth muscle cells and a chronic inflammation reaction. A variety of approaches are employed to lower the restenosis rates.
One solution is the use of biocorrodible metal alloys, because in most instances a permanent supporting function by the stent is not required; the initially damaged vessel is able to regenerate. In DE 197 31 021 A1 it is thus proposed, for example, to fabricate medical implants from a metal material, the major constituent thereof being an element selected from the group consisting of alkali metals, alkaline earth metals, zinc and aluminum. Alloys based on magnesium and zinc are described as being particularly suited. Minor constituents of the alloys can be manganese, cobalt, nickel, chromium, copper, cadmium, lead, tin, thorium, zirconium, silver, gold, palladium, platinum, silicon, calcium, lithium, aluminum, and zinc. Furthermore, DE 102 53 634 A1 discloses the use of a biocorrodible magnesium alloy comprising contents of magnesium >90%, yttrium 3.7 to 5.5%, rare earth metals 1.5 to 4.4% and the remainder <1%, which is suited in particular for producing an endoprosthesis, for example in the form of a self-expanding or balloon-expandable stent.
Biocorrodible implants thus constitute a promising approach to reducing the restenosis rate. One problem with implementing such systems is the corrosion behavior of the implant. For example, fragment formation due to the corrosion process should be suppressed to the extent possible until the implant has grown into the vessel well. In addition, the supporting function should be preserved for the therapeutically specified time frame.
Magnesium alloys are gaining increasing technical importance as materials, in particular also in medical technology, for example as materials for implants. So as to provide a brief overview of the different magnesium alloys, the classification and designation of alloys and the effects of the most important alloying elements will be described in more detail at this point. The designation of magnesium alloys according to the ASTM standard has become widely accepted around the globe. The designations of alloys consist of two letters of the major alloying elements, followed by the rounded contents thereof in percent by weight. The ASTM code designations for the alloying elements of magnesium assign the following letters to these alloying elements: A for aluminum; B for bismuth; C for copper; D for cadmium; E for rare earths; F for iron; H for thorium; K for zirconium; L for lithium; M for manganese; N for nickel; P for lead; Q for silver; R for chrome; S for silicon; T for tin; W for yttrium; Y for antimony; and Z for zinc. The code may be followed by suffix letters that indicate various development stages of the corresponding alloys (A, B, C . . . ). These letters generally designate the content of impurities—with the letter X designating the alloy as an experimental alloy. Example: According to the code designation, the alloy AZ91 D is a magnesium alloy with a nominal content of 9% by weight of aluminum and 1% by weight of zinc in the fourth development stage.
Because of the very promising properties of magnesium as a material for medical technology, attempts have been ongoing for quite some time to positively influence the properties profile thereof using suitable alloying elements. The increase in the mechanical properties of magnesium by alloying additions is based on solid solution hardening, precipitation hardening or grain size hardening. In addition to the mechanical properties, alloying additions can also be used to influence other important properties such as the corrosion resistance, castability and weldability, among other things.
The effects of the most important alloying elements of magnesium are summarized below in alphabetical order.
Al: Aluminum is the “classic” alloying element for magnesium. Aluminum is added to increase the tensile strength and hardness. In addition to increasing strength, aluminum causes a marked improvement in the castability of magnesium. The drawback is the increased tendency toward microporosity.
Ag: In conjunction with the rare earth metals, silver dramatically increases the temperature stability and creep resistance, but causes an increased tendency toward corrosion.
Be: Beryllium is added to the magnesium melt in extremely low concentrations (<30 ppm) in order to reduce the oxidation tendency of the melt.
Ca: Calcium has an effective grain refining effect and increases the creep resistance. However, when processed by way of casting, an increased tendency to adhere to the mold and increased hot cracking susceptibility are to be expected.
Mn: The most important effect of adding manganese is the strong improvement in corrosion resistance (the iron content is controlled by lowering the solubility).
RE: All rare earth metals (the element yttrium shall be included here) form eutectic phase diagrams with magnesium with limited solubility on the Mg-rich side, which allows precipitation hardening. Because very stable precipitations form, these elements dramatically increase the temperature stability and creep resistance of the alloys.
Si: The addition of silicon worsens the castability. Because very stable silicides (Mg2Si) can form, the creep resistance may be increased.
Zn: Like aluminum, zinc improves the castability and has a strength-increasing effect. However, as with aluminum, the tendency toward microporosity increases. Higher contents (>2%) increase the tendency toward hot cracking and worsen the weldability.
Zr: Zirconium is a very effective grain refining element. The fine grain size leads to an increase in the tensile strength, without lowering the strain. However, zirconium should not be added to melts containing aluminum or silicon because the reaction with these elements causes the grain-refining effect to be lost.
While magnesium alloys are used in many cases as cast alloys in lightweight construction (for example vehicle construction, aviation, mechanical engineering, consumer goods), what are known as wrought alloys are more customary in medical technology.
In non-ferrous metal metallurgy, wrought alloys refer to alloys that are distinguished from cast alloys by the ductility-favoring composition thereof, making them suitable for tasks such as rolling, pressing, drawing and forging, wherein the cast alloys are processed as liquid metal by pouring into a sand or permanent metal mold to form workpieces in the form of castings. Wrought alloys are an intermediate product, also referred to as a semi-finished product, the production of which is subject to some special characteristics as compared to cast alloys of the same type. In terms of the basic analytical composition, they differ only little from the cast alloys. Because the main requirement of a wrought alloy is the suitability for cold or hot forming, with the machinability optionally also being of importance, this may require that the content of some accompanying elements, which do not interfere with cast alloys, be limited and that elements that are suitable for promoting the further processing of the wrought alloy may need to be added.
The good properties of magnesium alloys are also offset by some negative aspects when using magnesium and the alloys thereof. Because magnesium crystallizes in the hexagonal close packing (hcp), the suitability for cold forming is poor in general. The reason for this is that below 225° C. deformation can only take place by two independent slip systems, and thus the condition for five independent slip systems (according to Von Mises) for a general change of shape is not met. Above approximately 225° C., the deformability increases almost suddenly because of the formation of new pyramidal slip planes—extensive deformations should thus occur above this temperature. Because of the hexagonal lattice structure and the tendency toward twinning, the magnesium material has thus become established as a wrought material only to a limited extent. The additional problem of the magnesium corroding requires the development of considerably more corrosion-resistant high-purity alloys with strict limits in terms of the contents of iron, nickel and copper (Fe<0.005%, Ni<0.001%, Cu<0.015%). For this reason as well, the available alloying range of magnesium as a wrought material is even further limited than for the cast alloys. Mg—Al alloys, such as AZ31, AZ61 and AZ80, play also a considerable role here, for example in lightweight construction. In addition, alloys containing zinc as the main alloying element, such as ZK60, exist. These alloys are processed by way of hot forming such as rolling, extrusion and forging at temperatures above 350° C. During a downstream cold forming operation, only small degrees of deformation are tolerated, otherwise cracks will form in the material. In the recent past, the development of wrought alloys of magnesium has been the focus of intensified interest again, because magnesium alloys are also to be used increasingly in medical technology workpieces, for example as implants, and more particularly as stent base bodies.
The prior art, for example DE 2008 10040143, describes magnesium alloys that contain yttrium and additional rare earth metals, because such an alloy stands out due to the physicochemical properties thereof and the high biocompatibility, in particular of the decomposition products thereof. Particularly preferred are magnesium alloys of the WE series, notably WE43, and magnesium alloys having a composition of 5.5 to 9.9% by weight of rare earth metals, of which yttrium can account for 0.9 to 5.5% by weight and the remainder ≦1% by weight, wherein the remainder can contain zirconium and/or silicon, and wherein magnesium accounts for the content in the alloy which is missing to make up 100% by weight. These magnesium alloys have already confirmed their particular suitability in experiments and initial clinical tests, which is to say they exhibit high biocompatibility, favorable processing properties, good mechanical characteristics and corrosion behavior that is adequate for the application purposes. In the present case, the collective term “rare earth metals” shall include scandium (21), yttrium (39), lanthanum (57) and the 14 elements following lanthanum (57), which is to say cerium (58), neodymium (60), promethium (61), samarium (62), europium (63), gadolinium (64), terbium (65), dysprosium (66), holmium (67), erbium (68), thulium (69), ytterbium (70) and lutetium (71).
Magnesium alloys for absorbable stents are often times provided with one or more coatings, for example a polymeric and/or optionally also active agent-containing coating, and the magnesium alloys must therefore have surface morphologies that exhibit sufficient adhesive action with respect to the coating, for example a polymeric primary coating (base coat). Otherwise the coating, and the polymer in particular, will tear open during dilation of the stent in the areas of the stents that have been deformed the most (for example on the insides of the arcs) and the metal will be exposed. The results are increased corrosion and a reduced supporting effect of the stent. This problem is not as pronounced in the WE 43 type magnesium alloys that have been used until now. The reason for this lies in the precipitations that are present due to the alloy, which in addition to the raised grain boundaries lead to a certain degree of surface roughness, even in electropolished surfaces. This in turn causes satisfactory adhesion of subsequent polymeric finishing coats.
Different novel and advantageous magnesium alloys, containing zinc as the essential major alloying element, for example magnesium alloys such as Z50 (95% by weight Mg; 5% by weight Zn) exhibit no tendency toward the formation of secondary phases and thus tend less toward precipitations due to the alloy. In addition, the grain boundaries thereof are not raised after electropolishing. In these magnesium alloys, containing zinc as the essential main alloying element, a considerable effect, which favors the polymer adhesion, is thus eliminated because of the very smooth surfaces following electropolishing.
Several experiments that improve the surface properties of implants, and of stents in particular, are already known from the prior art. US 2010/0305684, for example, discloses a Mg stent, which is coated with a ceramic layer so as to increase the corrosion resistance. By way of electrochemical fluorination, the Mg base body is provided with an intermediate MgF2 layer, which similarly to the final ceramic layer is intended to delay the corrosion of the Mg base body by preventing fractures and cracking. However, US 2010/0305684 does not contain any suggestions with regard to the surface characteristics of the intermediate MgF2 layer that is generated by way of electrochemical fluorination.
In addition, US 2006/0198869 discloses, in very general terms, the production of biodegradable stents. During the production, an etching step is carried out, and a coating made of polymer is also disclosed. The document also describes the creation of microstructures measuring 15 to 250 micrometers (m) in size by way of micro-etching. However, these microstructures according to US 2006/0198869 are formed only on the structural element that is additionally applied to the surface, in particular by and/or in the polymer material. In contrast, microstructuring of the surface of the Mg base body within the surface of the Mg alloy and/or the surface of the Mg alloy itself does not take place in US 2006/0198869.
The aim is thus to structure the surfaces of magnesium alloys, preferably containing zinc and/or aluminum, in particular containing zinc, as the essential main alloying element, in such a way that a roughness is achieved even in these magnesium alloys, which allows optimal adhesion of coatings, which is to say, for example, optimal polymer adhesion and/or optimal adhesion of an active agent-containing base coat in the case of a polymeric and/or optionally active agent-containing base coat.
The object according to the invention is thus in particular to structure the surfaces of such magnesium alloys, preferably containing zinc and/or aluminum, in particular zinc, as the essential main alloying element, for example magnesium alloys such as Z50, so that these surfaces exhibit increased adhesive strength for polymeric finishing coats and exhibit no drawbacks in terms of the breakage behavior. Another object is that of increasing the corrosion resistance of the base material, which is to say of the magnesium alloys, preferably containing zinc and/or aluminum, in particular zinc, as the essential main alloying element.